The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to the production of MRI perfusion images.
Any nucleus which possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus processes around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the magnetogyric constant gamma γ of the nucleus). Nuclei which exhibit this phenomena are referred to herein as “spins”.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net longitudinal magnetization M0 is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net longitudinal magnetization, M0, may be rotated, or “tipped” into the x-y plane to produce a net transverse magnetic moment Mt, which is rotating, or spinning, in the x-y plane at the Larmor frequency. The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the excitation signal B1 is terminated. There are a wide variety of measurement sequences in which this nuclear magnetic resonance (“NMR”) phenomena is exploited.
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (Gx, Gy, and Gz) which have the same direction as the polarizing field B0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified.
Perfusion as related to tissue refers to the exchange of oxygen, water and nutrients between blood and tissue. The measurement of tissue perfusion is important for the functional assessment of organ health. Images which show by their brightness the degree to which tissues are perfused can be used, for example, to assess the scope of brain tissues which have been damaged by a stroke, or to assess the scope of myocardial tissue damage resulting from a heart attack.
A number of methods have been used to produce perfusion images using magnetic resonance imaging techniques. One technique, as exemplified by U.S. Pat. No. 6,295,465, is to determine the wash-in or wash-out kinetics of contrast agents such as chelated gadolinium. In addition to the need for injection of a contrast agent, these methods require the acquisition and subtraction of baseline images.
Another class of MR perfusion imaging techniques attempts to measure blood flow by “tagging” or “labeling” spins flowing into a region of interest by applying RF excitation in an adjacent region and then acquiring image data from the region of interest. By subtracting a baseline image acquired without RF tagging, perfusion information is acquired and imaged. Repeated acquisitions and averaging of the results is used to improve perfusion image signal-to-noise ratio (SNR). Examples of these techniques are disclosed in U.S. Pat. Nos. 5,402,785; 6,285,900; 5,846,197; and 6,271,665 and the publications “Quantification Of Relative Cerebral Blood Flow Change By Flow-Sensitive Alternating Inversion Recovery Technique; Application to Functional Mapping” by S. G. Kim Magn. Reson. Med. 34(3):297-301, 1995; “MR Perfusion Studies With T1-Weighted Echo Planar Imaging”, by K. K. Wong et al Magn. Reson. Med. 34:878-887 (1995); and “QUIPSS II With Thin-Slice TI, Periodic Saturation” A Method For Improving Accuracy Of Quantitative Perfusion Imaging Using Pulsed Arterial Spin Labeling” by Luh et al Magn. Reson. Med. 41:1246-1254 (1999).
In all of these methods the amplitude or amplitude change of the NMR signal at each image voxel is the measure of perfusion at that location in the subject tissue. The basic structure of these NMR perfusion sequences includes one tagging slice and one imaging slice as shown in FIG. 3, separated by a distance (e.g., 5 mm) and excited at two different moments (e.g., 500 ms apart). If the tagging pulse inverts the magnetization by 180° in a tagging slice and there is flow of one cm/sec in the direction of the imaging slice, then the total magnetization M0 in this slice will be reduced when transverse magnetization is produce by an imaging pulse sequence. The detected NMR signal in a given voxel into which tagged spins flow will, therefore, be lower than without tagging. A similar effect can be obtained by pure saturation, i.e., by applying a tagging pulse flip angle equal to 90°. In this case, the signal reduction will be smaller. The levels of longitudinal spin magnetization M0 of inflowing tagged blood are shown in FIG. 4. Point Inv marks the longitudinal magnetization value for a 180° pulse, point Sat for a 90° pulse, and point Norm for a 0° tagging pulse. The general principle of flow detection is to subtract two images, one with no tagging and one which has been tagged. In the experiment illustrated in FIG. 3, only one flow velocity can be detected—exactly one cm/sec. Slower flowing blood will not arrive at the time of image acquisition; faster flowing blood will overshoot the slice. The sensitivity of this method is poor for several reasons: the T1 relaxation of blood is less than one second at a polarizing field of 3T, and the total volume of the microvascular structure is only a small part of the imaging voxel. To improve sensitivity, usually a plurality of imaging pairs is acquired and the differential signals are averaged. The repetition time (TR) has to be long enough for longitudinal magnetization to relax fully.